The present invention relates to the generation of magnetic fields, in particular for magnetic resonance imaging.
Magnetic Resonance Imaging (MRI) requires a high strength and uniformity magnetic field over a region of interest. In case of imaging of human subjects, it is also necessary that the shape of the magnet be suitable for easy positioning of the subject under examination, so that the volume to be examined is located in the region of homogeneous field, i.e. in the region where the magnetic field exhibits equal intensity and is unidirectional.
Electromagnetic fields are described by Maxwell's equations, which do not allow the existence of uniform fields in air, except in volumes completely enclosed within elements generating said fields. In general, however, the magnet cavity is designed to position the subject to be examined. Therefore the possibility of building a magnet able to generate a perfectly homogeneous field and simultaneously exhibiting openings suitable to position a subject is a priori excluded.
Therefore the art of manufacturing magnets suitable for MRI has been focused on finding techniques able to best approximate the required conditions, using magnetic field generators formed by conductive coils, carrying direct electric currents, or by blocks of magnetized materials. A number of techniques are used in the context, and they differ substantially if the field is generated by windings or by blocks of magnetized materials. From a fundamental point of view, anyway, the physical principles are the same.
The current state of the art in the field of MRI requires field intensities in the range from a minimum of 1 Tesla up to 8 or 9 Tesla. The most common structure of a magnetic field generator for MRI applications is a set of circular coils wound around different diameters and with different axial spacings along a single cylindrical symmetry axis. A proper choice of coil size and location enables the generation of fields having the features of intensity and homogeneity required over the volume of interest, which is located near the center of the structure. Moreover, the use of external coils wherein the current flows in opposite direction with respect to the internal coils enables “shielding”, i.e. reducing the field in regions where it is not necessary (or even harmful).
According to this structural choice, the magnet is shaped as a cylinder, and the subject is positioned inside said cylinder. The region of imaging is near the center of the cylinder, therefore it is hardly accessible from the outside. Since in the imaging region it is often positioned the head, or the chest, or the waist of the subject (for brain, or cardiac, or abdominal examinations respectively), the subject is enclosed in the cylindrical volume of the magnet, which leads to conditions of anxiety, discomfort or real claustrophobia, which make the examination impossible in a significant fraction of the cases.
Another structure that is employed is the “open” configuration in which an iron core comprising two circular or otherwise shaped pole pieces is excited by means of coils or of permanently magnetized material. This kind of magnets is limited in the maximum field intensity that can be reached either by the maximum remanence of the available permanent magnets or by the saturation characteristics of the iron constituting the core.
Still another “open” structure is based on the “split coil” or “Helmholtz coil” configuration in which two coaxial coils are spaced apart thus providing a space for introducing the patient in a direction orthogonal to the axis of the coils; in an alternate implementation the patient lies along the axis of the coils and the gap between them is used by a surgeon to access the patient body.
The design and manufacturing techniques for permanent magnets (i.e. based on the use of magnetized materials) are quite different. A permanent magnet can be manufactured by a combination of polyhedral blocks of a magnetized material, which has the property of generating a perfectly uniform field in a cavity inside the magnet when the cavity is completely enclosed. Removing a wall from the cavity causes a significant degradation of the field properties, which must be corrected by adding auxiliary blocks, suitably sized and positioned.
The use of permanent magnetic materials poses however some limits on the maximum achievable field intensity, which currently is about 0.5 to 0.8 Tesla, therefore below the levels currently provided by superconducting magnets.
As a consequence, the generation of a uniform magnetic field in a region of interest within a cavity or in a volume of open space, accessible to a patient, is generally achieved at the expense of efficiency, which is defined as the ratio between the energy of the magnetic field comprised in the spatial region of interest and the total energy used to generate said field, or otherwise as the ratio between the magnetic field intensity in the spatial region of interest and the peak current density flowing in the coils which generate said field.
According to the magnetic field theory, a region of uniform magnetic field in a volume free from electric currents (as is the region of interest for the applications here considered) can exist only around a saddle point. This is a consequence of the fact that maxima or minima of the field cannot exist in a volume of space free from electric currents. Moreover, the field intensity decreases when the distance from the generating currents increases.
The theoretical and practical problem of designing an open magnet therefore corresponds to the problem of determining a configuration suitable to generate a saddle point of the magnetic field in a region as far as possible from the magnetic structure, and therefore accessible, where however the field intensity is still sufficiently high for the desired application.
An overview of the known techniques to generate remote uniform magnetic fields, i.e. fields in spatial regions external to the field generating structure, is reported in the article “Generation of Remote Homogeneous Magnetic Fields” by Yuly M. Pulyer and M. I. Hrovat, published in IEEE Transactions on Magnetics, vol. 38, 2003 (1553). The paper describes the eleven configurations proposed in the literature to generate uniform magnetic fields for magnetic resonance imaging applications, both of open and closed type, and for each said configuration it reports the corresponding dipole-based model which describes schematically the main features of each structure. It is worth noting that all eleven models can be reduced to one or more dipoles of variable intensity, whose vectors are collinear, with either equal or opposite direction, or parallel, in which case they are never co-directionally oriented.
Typical examples of structures for the generation of a magnetic field with the desired properties are the configurations based on separated coils, which include two short solenoids facing each other, oriented either co- or contra-directionally, the configurations based on combinations of flat (“pancake”) coaxial coils, which generate different field intensities, or combinations of more complex structures having two or four parallel or antiparallel dipoles.
As a matter of fact, the structures so far practically used in MRI applications are the configuations based on separate coils, or on volume magnets shaped as C or H. U.S. Pat. Nos. 5,592,090 and 5,305,749 both disclose an open-structure magnetic assembly having a single source of magnetic field, in the form of a winding, and a structure of magnetic material for propagating the field excited by the source and establishing a uniform field in a region of a cavity intended for receiving a patient.
A deeper analysis of these structures allows their classification in terms of their “aperture degree”, which can be defined quantitatively by means of the aperture factor, defined as the ratio between the solid angle subtended by the total aperture, as seen from the center of the region of interest, and the total solid angle 4π.
It is obvious to a man skilled in the art that the greater is the number of open sides of the structure the lower is the efficiency of a magnetic assembly. For example, a solenoid or “tunnel” magnet is open at the opposite end faces, orthogonal to the direction of the magnetic field vector, while a C-shaped magnet is open at four faces parallel to the field (envelope of the magnet air gap), a “pancake” magnet including flat coaxial coils (as described e.g. in U.S. Pat. Nos. 4,701,736 or 5,428,292) is open at five faces, four parallel and one orthogonal to the field.
Let's consider for example a cylindrical coil (solenoid) with its length equal to its diameter, a C-shaped magnet with a pair of flat pole pieces with a circular cross section, separated by a distance equal to their diameter, and, by comparison, an infinite solenoid having the same diameter. The magnetic field in the region of interest is strictly related to the aperture factor. As a consequence, in a C-shaped magnet the ratio between distance and diameter of the pole pieces must be kept low, typically about 0.5, if a good efficiency is desired. Similarly, a solenoid should be at least 1.5 to 2 diameters long. The corresponding aperture factors would be 0.2929 for the solenoid and 0.7071 for the C-shaped magnet. However, while in a C-shaped magnet structure the field intensity in a median plane has a maximum near its center, in a solenoid structure the situation is opposite and the maximum for the field intensity is close to its inner walls.
From the point of view of efficiency the region of interest should be as close as possible to a wall rather than at the center of a current-free region.
These considerations are however conflicting with the requirements for particular MRI applications, e.g. for the study of the human motor cortex, where there is the need to keep the patient in a natural, erected or seated position, free to move.